The present invention relates to an image generation method in a scintillation camera, and in particular to method for interpreting the electrical signals received from an array of photomultiplier tubes.
In the human body, increased metabolic activity is associated with an increase in emitted radiation if the body is appropriately dosed with a radioactive tracer. In the field of nuclear medicine, increased metabolic activity within a patient is detected using a radiation detector such as a scintillation camera.
Scintillation cameras are well known in the art, and are used for medical diagnostics. A patient ingests, or inhales or is injected with a small quantity of a radioactive isotope. The radioactive isotope emits photons that are detected by a scintillation medium in the scintillation camera. The scintillation medium is commonly a sodium iodide crystal, BGO or other. The scintillation medium emits a small flash or scintillation of light, in response to stimulating radiation, such as from a patient. The intensity of the scintillation of light is proportional to the energy of the stimulating photon, such as a gamma photon. Note that the relationship between the intensity of the scintillation of light and the gamma photon is not linear.
A conventional scintillation camera such as a gamma camera includes a detector which converts into electrical signals gamma rays emitted from a patient after radioisotope has been administered to the patient. The detector includes a scintillator and photomultiplier tubes. The gamma rays are directed to the scintillator which absorbs the radiation and produces, in response, a very small flash of light. An array of photodetectors, which are placed in optical communication with the scintillation crystal, converts these flashes into electrical signals which are subsequently processed. The processing enables the camera to produce an image of the distribution of the radioisotope within the patient.
Gamma radiation is emitted in all directions and it is necessary to collimate the radiation before the radiation impinges on the crystal scintillator. This is accomplished by a collimator which is a sheet of absorbing material, usually lead, perforated by relatively narrow channels. The collimator is detachably secured to the detector head, allowing the collimator to be changed to enable the detector head to be used with the different energies of isotope to suit particular characteristics of the patient study. A collimator may vary considerably in weight to match the isotope or study type.
Scintillation cameras are used to take five basic types of pictures: spot views, whole body views, partial whole body views, SPECT views, and whole body SPECT views.
A spot view is an image of a part of a patient. The area of the spot view is less than or equal to the size of the field of view of the gamma camera. In order to be able to achieve a full range of spot views, a gamma camera must be positionable at any location relative to a patient.
One type of whole body view is a series of spot views fitted together such that the whole body of the patient may be viewed at one time. Another type of whole body view is a continuous scan of the whole body of the patient. A partial whole body view is simply a whole body view that covers only part of the body of the patient. In order to be able to achieve a whole body view, a gamma camera must be positionable at any location relative to a patient in an automated sequence of views.
The acronym xe2x80x9cSPECTxe2x80x9d stands for single photon emission computerized tomography. A SPECT view is a series of slice-like images of the patient. The slice-like images are often, but not necessarily, transversely oriented with respect to the patient. Each slice-like image is made up of multiple views taken at different angles around the patient, the data from the various views being combined to form the slice-like image. In order to be able to achieve a SPECT view, a scintillation camera must be rotatable around a patient, with the direction of the detector head of the scintillation camera pointing in a series of known and precise directions such that reprojection of the data can be accurately undertaken.
A whole body SPECT view is a series of parallel slice-like transverse images of a patient. Typically, a whole body SPECT view consists of sixty four spaced apart SPECT views. A whole body SPECT view results from the simultaneous generation of whole body and SPECT image data. In order to be able to achieve a whole body SPECT view, a scintillation camera must be rotatable around a patient, with the direction of the detector head of the scintillation camera pointing in a series of known and precise directions such that reprojection of the data can be accurately undertaken.
In generating an image with a nuclear scintillation camera, one of the problems encountered is that there is generally a shortage of detected gamma events.
One reason for the shortage of detected gamma events is that, for health reasons, a patient should be exposed to as little radiation as possible.
The image created by the scintillation camera is essentially a display of detected gamma events. If there are few counts, then there is little data to create the image, and the image may be meaningless from the point of view of human interpretation. It is not that the resolution is poor; it is just that the information is too sparse for a person to discern an image.
To generate an image from detected gamma events, the event information is written into an image or display matrix. Event by event, the data is written into picture elements or pixels. Each element or pixel contains input from zero to a high number of gamma events, proportional to the number of gamma events detected at the location corresponding to that pixel. The more gamma events, the brighter the pixel. A three dimensional graph of the pixels can be generated, showing the X and Y coordinates of the pixel locations in two dimensions, and the number of detected gamma events being indicated by the Z coordinate.
The collimator used in a scintillation camera provides the one to one spacial correlation of the emitted gamma rays at right angles to the crystal. The scintillation crystal used in nuclear scintillation cameras is sensitive. The collimator, however, reduces the efficiency greatly as gamma events occur in all directions, and as the collimator only lets through the gamma events that are substantially perpendicular to the scintillation crystal, most gamma rays are absorbed by the collimator. Collimators generally have efficiencies of minus four or five orders of magnitude; for example, for every 50,000 or so gamma events, only one passes through the collimator and is detected by the crystal.
Only a small amount of radioactive isotope can be administrated to the patient, and most of the gamma events go undetected. With so few counts, an image will not have enough information for form a recognizable picture. As more counts are detected, a pattern becomes discernable; however, details of the pattern cannot be made out; for example, the edge of an object will not be discernable.
Since the patient can only be exposed to a limited amount of radioactivity, one way to generate a better image is to take the picture, i.e. detect emitted gamma events, for a longer period of time. However, there is a limit to the length of time for which a patient can remain essentially motionless. And in some cases, it is impossible for the patient to remain motionless, such as when it is the patient""s heart that is being studied. It is common for studies to last for about twenty minutes, during which time the patient must attempt to remain as still as possible as any movement reduces the resolution of the generated image. As the study becomes longer, it becomes more difficult for a patient to remain still, and the resolution of the image tends to deteriorate.
One known method of dealing with the problem of a shortage of information is to apply a smoothing technique to the image data. Basically, smoothing techniques involve moving a certain amount of data from a pixel and moving it to surrounding pixels.
A typical technique or formula is a 121 242 121 smooth. The data associated with a particular pixel is assigned a weighting of 4 relative to its surrounding pixels. The surrounding orthogonal pixels are weighted as 2. The surrounding diagonal pixels are weighted as 1.
With respect to smoothing techniques, a heavy weighting means that the centre pixel is given an high weighting. An example would be a 121 2,20,2 121 smooth. A relatively small amount of data is assigned to surrounding pixels. This is referred to as a light smooth.
With basic smoothing techniques as discussed above, the data is moved without taking into account characteristics of the data as a whole; i.e. the same smoothing technique is applied to each pixel, without taking into account information from other pixels. The result is that the edges of the image become blurred.
A more sophisticated smoothing technique involves weighting the centre pixel by the median value of the nine pixels in the intermediate group. This is called a median smooth. The advantage is that one loses less resolution. The median smoothing technique was developed for looking at eye movements: since an eye generally looks quickly from one place to another.
In the preferred embodiment it is assumed that a tuning device exists, as described in U.S. Pat. No. 5,576,547 and U.S. Pat. No. 5,237,173 but not limited to such devices, and that the tuning is done before the acquisition for the energy information and positional information. The assumption is that before acquisition, tuning is performed on the detector head, which will normalize the responses of all the light detectors. The assumption is that the detector head is digital, but not limited to being digital. (This energy correction method can be used with any detector head on the market which can improve the characteristics of the detector heads.) After or instead of those tuning devices, a new calibration is also performed based on a hole phantom image acquisition.
Another smoothing technique examines the frequency content of the pixels. Smoothing is carried out in frequency space, or Fourier space. The resolution of the system (i.e. the camera that is writing the events into the pixels) can only resolve a certain spacial frequency and not higher. For example, with reference to the collimator, a camera may be able to resolve 4 mm line pairs (i.e. 2 mm of lead, 2 mm gap). This will give a frequency of 4 line pairs per cm. Any higher frequency than cannot be resolved. In between is statistical noise that does not really have a meaning. Thus, the frequency content in the pixels is examined. If the frequency content is above what the system can resolve, then the excess frequencies are filtered out.
Another smoothing technique uses a filter that implements a heavy smooth, and subtracts a light smooth and multiplied by a factor. Such a technique gives an edge enhancement that makes the image look better.
Smoothing techniques allow images to be discerned, but they do not add information. Such smoothing techniques simply spread out the known information so that information can be better interpreted by the human eye. However, in doing so, the spacial resolution of the image is compromised. In other words, the image looks better and patterns can be seen, but, in terms of information theory, information has actually been lost. It must be kept in mind that one will never be able to see something that cannot be seen from the raw or unsmoothed data.
To review, smoothing is generally required to create a recognizable image from insufficient data. However, resolution is lost during the smoothing process.
An object of the invention is to provide an improved image generation method in a scintillation camera.
A second object of the invention is to provide an improved method for interpreting the electrical signals received from an array of photomultiplier tubes.
According to the invention, there is provided a nuclear scintillation camera having a scintillation crystal for detecting a plurality of nuclear events and for generating a light scintillation corresponding to each detected nuclear event, an array of photodetectors for detecting light scintillations generated by the scintillation crystal, each light scintillation being detected by a plurality of the photodetectors in the array, each of the plurality of photodetectors generating an electrical signal corresponding to the intensity of light detected by that photodetector, a method for generating an image of the distribution and intensity of the nuclear events, the method comprising the steps of: (a) receiving signals from the plurality of photodetectors with respect to each nuclear event; (b) applying a plurality of positioning algorithms to the signals to calculate a plurality of position data, each position data being generated by each respective positioning algorithm; and (c) producing an image using the plurality of position data; whereby, when a small number of nuclear events are detected, a recognizable image can be obtained.
According to the invention, there is further provided a nuclear scintillation camera comprising: (a) a scintillation crystal for detecting a plurality of nuclear events and for generating a light scintillation corresponding to each detected nuclear event; (b) an array of photodetectors for detecting light scintillations generated by the scintillation crystal, each light scintillation being detected by a plurality of the photodetectors in the array, each of the plurality of photodetectors generating an electrical signal corresponding to the intensity of light detected by that photodetector; (c) means for receiving signals from the plurality of photodetectors with respect to each nuclear event; (d) means for applying a plurality of positioning algorithms to the signals to calculate a plurality of position data, each position data being generated by each respective positioning algorithm; and (e) means for producing an image using the plurality of position data; whereby, when a small number of nuclear events are detected, a recognizable image can be obtained.